Implantable neurostimulation systems have proven therapeutic in a wide variety of diseases and disorders. Pacemakers and Implantable Cardiac Defibrillators (ICDs) have proven highly effective in the treatment of a number of cardiac conditions (e.g., arrhythmias). Spinal Cord Stimulation (SCS) systems have long been accepted as a therapeutic modality for the treatment of chronic pain syndromes, and the application of tissue stimulation has begun to expand to additional applications such as angina pectoralis and incontinence. Deep Brain Stimulation (DBS) has also been applied therapeutically for well over a decade for the treatment of refractory chronic pain syndromes, and DBS has also recently been applied in additional areas such as movement disorders and epilepsy. Further, in recent investigations, Peripheral Nerve Stimulation (PNS) systems have demonstrated efficacy in the treatment of chronic pain syndromes and incontinence, and a number of additional applications are currently under investigation. Furthermore, Functional Electrical Stimulation (FES) systems, such as the Freehand system by NeuroControl (Cleveland, Ohio), have been applied to restore some functionality to paralyzed extremities in spinal cord injury patients.
Each of these implantable neurostimulation systems typically includes one or more electrode carrying stimulation leads, which are implanted at the desired stimulation site, and a neurostimulator implanted remotely from the stimulation site, but coupled to the stimulation lead(s). Thus, electrical pulses can be delivered from the neurostimulator to the stimulation lead(s) to stimulate or activate a volume of neural tissue. In particular, electrical energy delivered between at least one cathodic electrode and at least one anodic electrodes creates an electrical field, which when strong enough, depolarizes (or “stimulates”) the neurons beyond a threshold level, thereby inducing the firing of action potentials (APs) that propagate along the neural fibers.
Stimulation energy may be delivered to the electrodes during and after the lead placement process in order to verify that the electrodes are stimulating the target neural elements and to formulate the most effective stimulation regimen. The regimen will dictate which of the electrodes are sourcing current pulses (anodes) and which of the electrodes are sinking current pulses (cathodes) at any given time, as well as the magnitude and duration of the current pulses. The stimulation regimen will typically be one that provides stimulation energy to all of the target tissue that must be stimulated in order to provide the therapeutic benefit, yet minimizes the volume of non-target tissue that is stimulated. In the case of SCS, such a therapeutic benefit is “paresthesia,” i.e., a tingling sensation that is effected by the electrical stimuli applied through the electrodes.
While the electrical stimulation of neurons has generally been successful in providing a therapeutic benefit to the patient, there are instances where the target tissue is not directly adjacent to an electrode and, because the electrical field strength decreases exponentially with distance from the electrodes, a relatively strong electrical field must be created to generate APs in the target neural fibers. The electrical field may, however, also result in the generation of APs in the non-target neural fibers between the electrode and the target neural fibers. The generation of APs in the non-target neural fibers may, in turn, lead to undesirable outcomes (e.g., discomfort or involuntary movements) for the patient. Because the target neural tissue (i.e., the tissue associated with the therapeutic effects) and non-target neural tissue (i.e., the tissue associated with undesirable side effects) are often juxtaposed, therapeutically stimulating neural tissue while preventing side effects may be difficult to achieve. In the context of SCS, there may be a few ways of eliminating, or at least minimizing, the stimulation of non-target neural tissue.
For example, in the case where the electrode array is medio-laterally aligned (i.e., the electrodes are arranged transversely to the neural fibers of the spinal cord), it may be desirable to control the shape of the AP generating neural region of the spinal cord in order to prevent the generation of APs in non-target neural fibers. For example, to produce the feeling of paresthesia without inducing involuntary motor movements within the patient, it is often desirable to preferentially stimulate nerve fibers in the dorsal column (DC nerve fibers), which primarily include sensory nerve fibers, over nerve fibers in the dorsal roots (DR nerve fibers), which include both sensory nerve fibers and motor reflex nerve fibers. While DC nerve fibers are the intended targets in conventional SCS, in fact, the DR nerve fibers often are recruited first because of geometric, anatomical, and physiological reasons. For example, the DR nerve fibers have larger diameters than the largest nearby DC nerve fibers, and thus, have a lower threshold at which they are excited. Other factors that contribute to the lower threshold needed to excite DR nerve fibers are the different orientations of the DC nerve fibers and DR nerve fibers, the curved shape of the DR nerve fibers, and the inhomogeneity and anisotropy of the surrounding medium at the entrance of the DR nerve fibers into the spinal cord. Thus, DR nerve fibers may still generate APs at lower voltages than will nearby DC nerve fibers. As a result, the DC nerve fibers that are desired to be stimulated have a lower probability to be stimulated than do the DR nerve fibers, and thus, the reflex motor nerve fibers intermingled among the sensor nerve fibers of a dorsal root are often recruited, leading to discomfort or muscle twitching, thereby preventing satisfactory paresthesia coverage.
For reasons such as these, it is often desirable to modify the threshold at which neural tissue is activated in a manner that maximizes excitation of the target neural tissue, while minimizing excitation of the non-target neural tissue; that is, to increase the DR/DC threshold ratio. This can be accomplished by sinking an electrical pulse to a cathodic electrode located at the center of the spinal cord to depolarize the target tissue adjacent the cathodic electrode, thereby creating APs along the DC nerve fibers, while an electrical pulse can be sourced to anodic electrodes on both sides of the cathodic electrode to hyperpolarize non-target tissue adjacent the anodic electrodes, thereby increasing the threshold of the DR nerve fibers.
As another example, in the case where the electrode array is rostro-caudally aligned (i.e., the electrodes are arranged along the neural fibers of the spinal cord), it may be desirable to induce APs in a bundle of target nerve fibers, and to the extent that APs are induced in bundle of non-target nerve fibers, block APs within the non-target nerve fibers from reaching the brain or any other parts of the nervous system. In particular, an electrical pulse can be sunk to a cathodic electrode to depolarize target tissue adjacent the cathodic electrode, thereby creating APs along a first bundle of nerve fibers, while an electrical pulse can be sourced to one or more anodic electrodes above or below the cathodic electrode to hyperpolarize non-target tissue adjacent the anodic electrode(s), thereby blocking any APs along a second bundle of nerve fibers that were inadvertently induced by the sink electrical pulse of the cathodic electrode.
Because the amount of electrical current that is sourced must equal the amount of electrical current that is sunk, the amount of sourced electrical current must be limited in order to minimize the adverse effects that could potentially occur as a result of the increased amount of the sunk electrical current. For example, in the previously described case where the electrode array is rostro-caudally aligned, an increase in the electrical current sunk by the cathode as a result of an increase in the electrical current sourced by the anodes(s) may result in the generation of APs in non-target nerve fibers that are not blocked by the sourced electrical current. In the previously described case where the electrode array is medio-laterally aligned, an increase in the electrical current sunk by the cathode as a result of an increase in the electrical current sourced by the anodes may result in the generation of APs in non-target DC nerve fibers.
To limit the amount of current sunk by a cathode, it is known to redistribute some of the cathodic current to a large surface area, such as the case of the IPG. Such a technique is described in U.S. patent application Ser. No. 11/300,963, entitled “Apparatus and Methods for Stimulating Tissue,” which is expressly incorporated herein by reference. By distributing the cathodic current to a surface area that has no, or very little, effect on the neural tissue, the magnitude of the electrical pulses sourced by the anodes can be increased while avoiding a commensurate increase in the magnitude of the electrical pulses sunk to the cathode that is adjacent the neural tissue. In this manner, any adverse effects that may otherwise occur as a result of an increase in the electrical current sunk to the cathodic electrode, and thus delivered through the neural tissue adjacent the cathodic electrode, can be minimized.
While this electrical current redistribution technique is beneficial, it can only be implemented within an IPG that has independent current or voltage sources for the electrodes. That is, an IPG with a single current or voltage source provides no means for redistributing a selected amount of cathode current to the IPG case. Furthermore, inadvertent stimulation of tissue in the pocket in which the neurostimulator is implanted may occur. This pocket stimulation problem is exacerbated when a microstimulator, which is an implantable neurostimulator in which the body or case of the device is compact (typically on the order of a few millimeters is diameter by several millimeters to a few centimeters in length), is used to deliver energy to the stimulation lead. Because the case of a microstimulator is relatively small, the current density on the surface of the case may be quite high when the microstimulator is operated in a monopolar mode. As a result, undesired and perhaps annoying or painful stimulation in the implantation pocket might be expected.
There, thus, remains a need for an alternative neurostimulation method and system that minimizes any adverse effects that may result in an increase in cathodic current when the anodic current is increased.